Two-pole impedance sensors have been proposed for various fields of sensor technology. Such two-pole impedance sensors are described in WO 93/22678, DE 19610115 A1, U.S. Patent Ser. No. 60/007840, and Peter Van Gerwen et al., “Nanoscaled Interdigitated Electrode Arrays for Biochemical Sensors”, Proc. International Conference on Solid-State Sensors and Actuators (Transducers '97), pp. 907-910, 1997, for the area of biosensor technology, and Hagleitner, C. et al. “A Gas Detection System on a Single CMOS-Chip comprising capacity, calorimetric, and mass sensitive microsensors”, Proc. International Solid-State Circuit Conference (ISSCC), p. 430, 2002, proposes a two-pole impedance measurement for the sensor technology of other chemical substances, for example gas sensors.
A description is given below, referring to FIG. 1A to FIG. 5B, of a sensor arrangement in accordance with the prior art which represents a DNA sensor on the basis of the two-pole impedance method.
FIG. 1A, FIG. 1B show an interdigital electrode arrangement 100, in which a first electrode structure 102 and a second electrode structure 103 are applied in a substrate 101, said electrode structures clearly being interdigitated. FIG. 1A shows a plan view of the interdigital electrode arrangement 100, and FIG. 1B shows a cross-sectional view along the section line I-I′ shown in FIG. 1A. The interdigital electrode arrangement 100 contains periodic electrode components of the electrode structures 102, 103 arranged next to one another.
The structure shown in FIG. 1A, FIG. 1B comprises electrodes arranged periodically next to one another, so-called interdigital electrodes.
In order to explain the principle of the functioning of the interdigital electrode arrangement 100, a first partial region 104 of the interdigital electrode arrangement 100 is described with reference to FIG. 2A, FIG. 2B.
The first partial region 104 is shown as a cross-sectional view in a first operating state in FIG. 2A, and as a cross-sectional view in a second operating state in FIG. 2B.
Capture molecules 200 (DNA half strands) are in each case immobilized on the electrode structures 102, 103. Gold material is preferably used for the electrode structures 102, 103, so that the immobilization of the capture molecules 200 is realized using the particularly advantageous gold-sulfur coupling known from biochemistry, by virtue of, by way of example, a thiol group (SH group) of the capture molecules 200 being chemically coupled to the gold electrodes 102, 103.
Situated above the sensor electrodes 102, 103 during active sensor operation is an electrolytic analyte 201 to be examined, which is intended to be examined for the presence of particles 202 to be detected (for example specific DNA molecules). A hybridization, that is to say a binding of DNA strands 202 to the capture molecules 200, is effected only when the capture molecules 200 and the DNA strands 202 match one another in accordance with the key-lock principle (cf. FIG. 2B). If this is not the case, then hybridization is not effected (not shown). The specificity of the sensor is thus derived from the specificity of the capture molecules 200.
The electrical parameter evaluated during this measurement is the impedance Z 203 between the electrodes 102, 103, which is illustrated schematically in FIG. 2A, FIG. 2B. On account of a hybridization that has taken place, the value of the impedance changes since the DNA particles 202 to be detected and the capture molecules 200 are composed of a material which has electrical properties deviating from those of the material of the electrolyte, and, after the hybridization, the electrolyte is clearly displaced from the volume region surrounding the electrodes 102, 103.
FIG. 3 shows a second partial region 105 (cf FIG. 1B) of the interdigital electrode arrangement 100 in a cross-sectional view.
The second partial region 105 represents a larger partial region of the interdigital electrode arrangement 100 than the first partial region 104 illustrated in FIG. 2A, FIG. 2B. FIG. 5 schematically shows the profile of the electric field lines 300 between respectively adjacent electrode structures 102, 103. As is furthermore shown in FIG. 3, the field profiles are periodic within a respective region imagined by virtue of two lines of symmetry 501, with the result that the consideration of two directly adjacent electrode structures 102, 103 as shown in FIG. 2A, FIG. 2B is sufficient. Furthermore, a coverage region 302 is shown schematically in FIG. 3 for each of said electrode structures 102, 103, said coverage region representing the capture molecules 200 immobilized on the electrode structures 102, 103 and particles 202 to be detected that have possibly hybridized with said capture molecules. It can clearly be understood from the illustration shown in FIG. 3 that the profile of the field lines 300 is significantly influenced on account of a hybridization event since the physical-chemical properties of, in particular, the coverage region 302 are altered.
It should furthermore be noted that capture molecules may supplementarily or alternatively be provided in regions between electrodes 102, 103 (not shown). The electrical properties of the electrolyte once again change in the case of hybridization events between particles to be detected and capture molecules provided in regions between the electrodes.
FIG. 4 schematically shows a simplified equivalent circuit diagram 400 of the first partial region 104 of the interdigital electrode arrangement 100 as shown in FIG. 2A. The equivalent circuit diagram 400 shows a variable first capacitance 401 CM, the value of which is dependent on the extent of a hybridization that has taken place at the electrode structure 102. A variable first nonreactive resistance 402 RM is connected in parallel therewith. The components 401, 402 clearly represent the electrical properties of the region surrounding the first electrode structure 102. Furthermore, a variable second capacitance 403 CE and a variable second nonreactive resistance 404 RE connected in parallel therewith are shown, which represent the electrical properties of the analyte 401. Also shown are a variable third capacitance 405 CM and a variable third nonreactive resistance 406 RM connected in parallel therewith, representing the electrical properties of the region surrounding the second electrode structure 103. As is furthermore shown in FIG. 4, the parallel circuit comprising components 401, 402, the parallel circuit comprising components 403, 404 and the parallel circuit comprising components 405, 406 are connected in series. The components 401 to 406 are represented as variable in order to illustrate that their values may change on account of a sensor event.
As is shown in the equivalent circuit diagram 500 of the first partial region 104 as shown in FIG. 5A, an AC voltage V is applied to one of the electrodes 102, 103 in order to determine the value of the impedance. The AC voltage V is provided using an AC voltage source 502. The AC current I flowing through the arrangement is detected using the ammeter 501. The components 501, 502 are connected in series with one another and are connected between the parallel circuit comprising components 405, 406 and the electrical ground potential 503. The AC current signal I resulting at the electrodes 102, 103 is evaluated together with the applied AC voltage V in order to determine the impedance. As an alternative, a signal, that is to say an electrical voltage, may in each case also be applied to both electrodes 102, 103; the signals are then in antiphase.
The version of a simplified equivalent circuit diagram 510 as shown in FIG. 5B differs from the equivalent circuit diagram 500 shown in FIG. 5A in that the elements CM 401, 405 and RM 402, 406 are combined to form a first effective capacitance 2CM 511 and, respectively, to form a first effective nonreactive resistance 512 2RM.
The distance between the electrodes 102, 103 typically lies in the sub-μm range. In accordance with the interdigital electrode arrangement 100, a multiplicity of electrode components (clearly fingers) of the electrode structures 102 and 103 are arranged parallel. Circular arrangements can be used for reasons of fluidics, as is described in R. Thewes et al., “Sensor Arrays for Fully Electronic DNA Detection on CMOS”, Proc. Int. Solid-State Circuits Conf. (ISSCC), p. 350, 2002, for a different detection method based on the use of interdigital electrodes. The external dimensions or the diameter of such individual sensors is in the region of approximately 100 μm or even less down into the single-digit mm range.
With regard to the exciting AC voltage V, it should be taken into consideration that its root-mean square value or its peak value is not to exceed a specific maximum value. The biochemical or electrochemical boundary conditions that enable the operation of such sensors are violated when such a maximum value is exceeded. If the electrode potential (which is referred to the electrical potential of the electrolyte) exceeds an upper threshold value, then specific substances in a region surrounding an electrode may be oxidized. If the electrical potential (which is referred to the electrical potential of the electrolyte) falls below a lower threshold value, substances are reduced there. An undesirable oxidation or reduction may have the effect, inter alia, that the chemical bonds entered into in the course of immobilization and hybridization are broken. Furthermore, electrolysis may commence at the sensor electrodes, with the result that the electrolysis products bring the chemical milieu required for operation of the sensors out of the required equilibrium and lead to gas formation. The absolute values of the critical potentials depend on the composition and the concentration ratio of the chemical surroundings of the electrodes (for example immobilization layer, analyte, etc.).
Typical values for the exciting voltage lie in the range of a few 10 mV to at most in the region around 100 mV. This is an important boundary condition for the operation of such sensors since the resulting measurement signal (current intensity I), with regard to its magnitude, is approximately directly proportional to the applied voltage.
There is often an interest in carrying out not just one test on a sensor arrangement but many tests on a suitable sample, the analyte, temporally in parallel. Miniaturized bio-/chemosensor arrays that can be realized on corresponding chips serve for parallel detection of different molecules or different substances in an analyte to be examined. A large number of the corresponding electrical sensors are realized on a glass, plastic, silicon or other substrate. On account of the property of parallelization, such sensor array chips including a corresponding evaluation system are afforded diverse applications in medical diagnosis technology, in the pharmacological industry (for example in the area of pharmacological screening, “high throughput screening”, HTS), in the chemical industry, in foodstuffs analysis, in ecological and foodstuffs technology and analysis, etc.
However, impedance sensors (such as the one described with reference to FIG. 1A to FIG. 5B) have hitherto been presented only as individual sensors or in small arrays, in principle comprising a stringing together of individual sensors.
In order to be able to carry out a large number of tests on an analyte temporally in parallel, it is endeavored to arrange a larger number of such sensors specified with respect to different substances in an array on a chip. Realizing arrays with two-pole impedance sensors gives rise to the challenge that the signals of all the sensors have to be fed to a read-out device. If e.g. a passive chip with 8×12=96, 32×48=1536 or generally m×n positions is present, 2×96=192, 2×1536=3072 or 2×m×n individual pads are present. Each sensor has to be separately readable, and the number of chip pads used is not to be too high on account of the production outlay for chip and reader and primarily for reasons of security in contact-connection (signal integrity). The simple approach, e.g. connecting all pads to the reader, yields 2×m×n (that is to say, in the example, 192 or 3052) pads. This is considerably too large for practical applications. This similarly holds true for the approach of operating one electrode in a common fashion and connecting all the remaining electrode terminals and also the common electrode to the reader. In this case, the number of pads is admittedly lower (n×m+1, that is to say 97 or 1537 in the examples), but is still very large.
One possibility is to use so-called active chips which, apart from the materials for the transducers (in particular the sensor electrodes, e.g. gold for the interdigital electrodes), provide additional active circuits for the signal preprocessing and the multiplexing of signals on-chip and also corresponding wiring planes. A solution of this type is described in R. Thewes et al. for a different method that is likewise based on the use of interdigital electrodes.
The prior art discloses biosensors that operate in accordance with the principle of redox recycling (cf. R. Hintsche et al, “Microelectrode arrays and application to biosensing devices”, Biosensors & Bioelectronics, pp. 697-705, 1994, R. Hintsche et al., “Microbiosensors using electrodes made in Si-technology” in “Frontiers in Biosensorics I”, F. Scheller et al. ed., Birkhauser, Basel, Switzerland, 1997, M. Paeschke et at, “Highly sensitive electrochemical microsensors using submicrometer electrode arrays”, Sensor and Actuators B, pp. 394-397, 1995, and WO 00/62048).
FIG. 1C, which illustrates an interdigital electrode arrangement as in FIG. 1A, FIG. 1B, shows a redox recycling biosensor 110 in accordance with the prior art.
Such a redox recycling sensor 110 requires a four-electrode system formed on a substrate 111. The redox recycling biosensor 110 has a first interdigital electrode 112 and a second interdigital electrode 113 that is clearly intermeshed with the latter in a finger-type fashion. Typical values for the width of and the distance between the interdigital electrodes 112, 113 lie in the range of between approximately 0.5 μm and approximately 2 μm.
Capture molecules (not shown) are immobilized on the interdigital electrodes 112, 113, and can hybridize with particles to be detected in a solution to be examined. In accordance with the redox recycling principle, during operation of the redox recycling biosensor 110, oxidation and reduction processes of electrochemically activated particles are effected in the case of a sensor event at the interdigital electrodes 112, 113. By means of a label molecule that is chemically bound to particles of an analyte that are to be detected and have hybridized with capture molecules, electrochemically activated particles are produced if a special component is added to the solution.
The further electrodes shown in FIG. 1, namely a reference electrode 114 and a counterelectrode 115, form, together with a differential amplifier 116, a potentiostat. A noninverting input 116a of the differential amplifier 116 is coupled to the electrical ground potential 117, whereas an inverting input 116b of the differential amplifier 116 is coupled to the reference electrode 114. An output 116c of the differential amplifier 116 is coupled to the counterelectrode 115. The electrochemical potential of an electrolyte introduced into the redox recycling biosensor 110 is measured by means of the reference electrode 114. By means of the counterelectrode 115, the differential amplifier 116 is regulated as a regulating amplifier to a predetermined electrical potential. It should be noted that the predetermined potential is clearly the electrical potential at the noninverting input 116a of the differential amplifier 116, the electrical ground potential 117 in accordance with the example shown.
For the operation of the redox recycling biosensor 110, electrical voltages having a different sign relative to the reference potential (ground potential 117) are applied to the interdigital electrodes 112, 113, which are operated as generator and collector electrodes, by means of a first DC voltage source 118 and by means of a second DC voltage source 119. The magnitude of said voltages is to be chosen to be high enough that the processes of oxidation and reduction of the redox recycling substance function reliably and efficiently, but on the other hand they are not to be chosen too high in order to avoid unintended electrochemical processes (e.g. electrolysis) at the electrodes 112, 113.
It should be noted that very precise threshold values for the commencement of oxidation and reduction of specific substances do not exist. In practice, values are chosen whose magnitude in each case lies sufficiently reliably (typically a few 10 mV) above these threshold ranges. In practice, the applied voltages lie in the region of a few 100 mV (for example +300 mV and −100 mV) for the substances used in R. Thewes et al.
As is furthermore shown in FIG. 1C, a first ammeter 120 coupled to the first DC voltage source 118 and a second ammeter 121 coupled to the second DC voltage source 119 are provided. The first ammeter 120 can be used to detect an electric current resulting from the applied voltages and the reduction/oxidation processes at the second interdigital electrode 113. The second ammeter 121 can be used to detect an electric current resulting from the applied voltages and the reduction/oxidation processes at the first interdigital electrode 112. The values of said currents are a characteristic measure of the sensor events that have taken place at the electrodes 112, 113 and thus of the concentration of particles to be detected in a solution to be examined.
Electrochemical sensors known from the prior art, such as the redox recycling biosensor 110, have hitherto been presented only as individual sensors in small “passive” arrays, clearly as a stringing together of a few individual sensors, or in “active” arrays which, apart from the actual biosensors, contain further active components on which either switching matrices, cf. WO 00/62048, or complex circuit technology, cf. R. Thewes et al., are or is contained.
For many applications, it is desirable to carry out not just one test with a biosensor but many tests in a given sample, the analyte, temporally in parallel. Miniaturized bio-/chemosensor arrays that can be realized on corresponding chips serve for the temporally parallel detection of different particles to be detected in an analyte to be examined. A large number of the corresponding sensors can be realized on a substrate, for example on a glass, plastic, silicon or other substrate. On account of the high degree of parallelization, such sensor arrangements including a corresponding evaluation system are afforded diverse applications in medical diagnosis technology, in the pharmacological industry (e.g. for pharmacological screening, “high throughput screening”, HTS), in the chemical industry in foodstuffs analysis, in ecological and foodstuffs technology and analysis.
Realizing passive arrays with sensors in accordance with the principle of redox recycling gives rise to the challenge that the signals of all the sensors have to be fed to a read-out device. If, by way of example, a passive substrate with 8×12=96, 32×48=1536 or generally in the case of a matrix-type arrangement with m rows and n columns m×n positions is present, this requires 2×96=192, 2×1536=3072 or generally 2 m×n separate electrode terminals of the sensors and also two further terminals for the reference electrode and counterelectrode of the potentiostat arrangement. Each biosensor zone should be separately readable, and the number of chip terminals (“pads”) used is not to be too high for reasons of the outlay (for chip and read-out device) and primarily for reasons of security in contact-connection (signal integrity).
A simple approach, for example coupling all the electrode terminals to the read-out device, yields 2 m×n+2, that is to say, in the example, 194 or 3074, terminals and therefore cannot be realized or cannot be realized satisfactorily. This similarly holds true for the approach of operating one electrode of all the sensor zones in a common fashion and coupling all the remaining electrode terminals and also the common electrode to the read-out device. In this case, the number of terminals is admittedly lower, namely n×m+1+2 (99 or 1539 in the example), but still much too large.
The disadvantage of the approaches described in WO 00/62048, R. Thewes et al., which make use of active arrays with a complicated CMOS technology for the individual sensor zones, consists in the high manufacturing costs compared with purely passive arrays in which CMOS circuits are obviated for the sensor zones.
WO 96/07917 discloses a self-addressable, self-assembling microelectronic system for carrying out molecular diagnosis, analysis, multiple step and multiplex reactions in microscopic formats.
WO 01/43870 A2 discloses a method and an apparatus which have a platform for a column- and row-addressable high-density biosensor arrangement.
WO 01/75151 A2 discloses a method for detecting macromolecular biopolymers by means of an electrode arrangement.
WO 88/09499 A1 discloses an optimized capacitive sensor for chemical analysis and measurement.